Spectroscopic analysis apparatus and method with excitation system and focus monitoring system

ABSTRACT

The present invention relates to an analysis apparatus, in particular a spectroscopic analysis apparatus, for analyzing an object, such as blood of a patient, and a corresponding analysis method. To aim the confocal detection volume inside a blood vessel orthogonal polarized spectral imaging (OPSI) is used to locate blood capillaries in the skin. An imaging system (img) with slightly shifted imaging planes (i 1 , i 2 ) for OPS imaging of blood vessels is proposed to provide auto-focusing. The confocal Raman detection plane (dp) is located in between these two imaging planes (i 1 , i 2 ). Based on measured amount of defocus for the imaging planes (i 1 , i 2 ), the focusing of the imaging system (img), the excitation system (exs) for exciting the target region and the detection system (dsy) are adjusted such that the difference of the amount of defocus equals a predetermined amount so that the confocal detection plane (dp) is located inside the blood vessel (V). Thus, continuous auto-focusing with high accuracy can be achieved. The invention relates also to an optical tracking system for continuously tracking a point of a moving object (obj).

The present invention relates to an analysis apparatus, in particular aspectroscopic analysis apparatus, for analyzing an object, such as theblood of a patient, and a corresponding analysis method. Further, theinvention relates to an optical tracking system for continuouslytracking a point of a moving object.

In general, analysis apparatuses, such as spectroscopic analysisapparatuses, are used to investigate the composition of an object to beexamined, e.g. to measure the concentration of various analytes in bloodin vivo. In particular, analysis apparatuses employ an analysis, such asa spectroscopic decomposition, based on interaction of the matter of theobject with incident electromagnetic radiation, such as visible light,infrared or ultraviolet radiation.

A spectroscopic analysis apparatus comprising an excitation system and amonitoring system is known from WO 02/057759 A2 which is incorporatedherein by reference. The excitation system emits an excitation beam toexcite a target region during an excitation period. The monitoringsystem emits a monitoring beam to image the target region during amonitoring period. The excitation period and the monitoring periodsubstantially overlap. Hence the target region is imaged together withthe excitation, and an image is formed displaying both the target regionand the excitation area. On the basis of this image, the excitation beamcan be very accurately aimed at the target region.

The analysis method known from WO 02/057759 A2 for simultaneous imagingand spectral analysis of a local composition is done by separate lasersfor confocal video imaging and Raman excitation or by use of a singlelaser for combined imaging and Raman analysis. Orthogonal polarizedspectral imaging (OPSI), which is also described in WO 02/057759 A2, isa simple, inexpensive and robust method to visualize blood vessels closeto the surface of organs which can also be used to visualize bloodcapillaries in the human skin. Blood capillaries close to the skinsurface have a diameter of about 10 μm. Due to confocal detection thesource of collected Raman signals is well confined in all threedimensions inside a spot of a size smaller than 5×5×5 μm³. This allowscollecting Raman signals from blood without background signal from skintissue if the focus is located in a blood capillary. This spot locationis possible if the lateral position of the blood vessel as well as thedepth of the vessels below the skin surface are known with a resolutionof 1 μm or better.

Because of an effective back-illumination of blood vessels, OPSI isessentially a 2-dimensional technique. The only depth information isobtained by the influence of the amount of (de)focus on the images. Ifan objective with a numerical aperture (NA) higher than 0.8 is used, thedepth of field in skin is below 0.5 μm. Therefore, with accuratefocusing algorithms based on image analysis it is possible to find thedepth of the blood vessel.

Known auto-focusing methods are based on scanning the axial position ofthe objective focusing the imaging beam and the confocal excitation beamonto the object of interest, while measuring the value of a meritfunction to quantify the amount of (de)focus. The best focus is found byoptimizing the value of the merit function. In general there are manypossibilities to change the focus position. For instance, one or twolenses in the objective can be moved (as in a photo camera) or anotherlens in the system can be moved. Also the shape of an optical element inthe system can be changed, for example an electro wetting opticalelement. However, if the object is not known, the maximum of the meritfunction is also unknown. Therefore, the merit function provides onlyinformation about the amount of focus in relation to other focuspositions.

Patients, however, will move in lateral as well as in transversaldirections. Therefore, continuous measurement and adjustment of theoptimal location of the confocal detection center is required.Transversal movements in the image plane can easily be detected, whereasaxial movements (perpendicular to the detection plane) are much moredifficult to detect. A common method of detecting axial movement ordefocus is by continuously moving the detection plane around the centralbest focus position (so called wobbling). This can be done by moving theimaging objective or another optical element in the imaging system. Ifthe focus becomes better in front or behind the central position, thecentral position of the objective is changed. In known systems thedetection volume is located in the image plane. Therefore this detectionvolume is also continuously moved around the best measurement position.This has the disadvantage that the confocal detection volume is locatedinside a blood vessel for only a fraction of time, and to avoid mixingof skin spectra with blood spectra, the intake of Raman signal has to begated. This increases the time needed to collect sufficient Ramansignal, which is in case of continuous recording already at least 30sec.

Further disadvantages are, that due to changes in the blood flow theshape and size of a capillary change continuously; so that comparingimages acquired at different times add uncertainty to the position ofbest focus. Additionally, the fact that more time is needed to collectsufficient Raman signal adds to the noise in the Raman spectrum becausemore dark current is acquired or because more readout noise is added.

It is therefore an object of the present invention to provide anoptimized analysis apparatus and a corresponding analysis method forimaging and spectroscopic analysis of an object which allow continuousaccurate auto-focusing of the excitation beam onto the object, inparticular a blood vessel, even during movements of the object withoutchanging the position of the detection volume continuously.

This object is achieved according to the present invention by ananalysis apparatus as claimed in claim 1 comprising:

an excitation system for emitting an excitation beam to excite a targetregion,

a monitoring system comprising a monitoring beam source for emitting amonitoring beam and an imaging system to image the target region,

a detection system for detecting scattered radiation from the targetregion generated by the excitation beam,

focusing means for focusing the imaging system on at least twosubstantially parallel imaging planes at a predetermined distance andfor focusing the detection system and the excitation system on adetection plane substantially parallel and in-between the imaging planesat predetermined distances,

defocus detection means for determining the amount of defocus of theimaging system from the detection plane for the at least two imagingplanes, and

auto-focusing means for controlling the focusing means to commonlychange the focusing of the imaging system, the excitation system and thedetection system based on the determined amount of defocus and thepredetermined distances between the imaging planes and the detectionplane such that the difference between the determined amount of defocusfor the at least two imaging planes equals a predetermined amount.

The object is further solved by a corresponding analysis method asclaimed in claim 10. Preferred embodiments of the invention are definedin the dependent claims.

The invention is based on the idea to take at least two images of theobject of interest (e.g. a blood vessel) at different imaging planes(focusing planes), i.e. the imaging planes of the imaging system areslightly separated in axial direction (direction of the monitoringbeam). The confocal Raman detection center, i.e. the detection planeonto which the excitation beam emitted by the excitation system forexciting the target region (the blood vessel) and the detection systemare focused, is located in between these two imaging planes whereby thedistances between the imaging planes and the detection plane arepredetermined and known. The amount of defocus of the imaging system,i.e. of the at least two images acquired at the imaging planes, isdetermined with a merit function by the defocus detection means. Sincethe size and shape of the object of interest (e.g. blood vessels) arenot known the merit function provides only a relative amount of defocus.This (relative) amount is then used for control of the focusing means toadjust the position of the detection plane and the imaging planes incommon such that the difference between the determined amount of defocusfor the at least two imaging planes equals a predetermined amount forthe at least two imaging planes. With this information the Ramanconfocal detection volume can be continuously located exactly inside theobject of interest (such as a blood vessel).

Compared to other known auto-focusing techniques the present inventionprovides the advantage that the confocal detection volume iscontinuously located in the center of the object of interest, even ifthe object moves during the measurement. According to preferredembodiments no moving elements are required and a single microscopeobjective, having a high numerical aperture can be used as focusingmeans. No continuous movement of the detection plane around the centralbest focus position (wobbling) is required.

According to a preferred embodiment exactly two images are acquiredbeing located on different sides of the detection plane, preferably atequal distances. The control of the focusing means is therein adaptedsuch that the amount of defocus for each imaging plane is substantiallyequal, eventually with a certain offset. Thus, by the merit function itis checked how much focused the image for each imaging plane is, and thefocusing means is adapted such that the amount of defocusing for eachimaging plane becomes substantially equal. If the patient moves duringthe analysis, this movement is continuously detected so that thefocusing means are continuously readjusted automatically.

According to another preferred embodiment the imaging system comprisestwo cameras, for instance CCD-cameras or CMOS cameras, each beingfocused on one of the imaging planes. Both cameras simultaneously takean image of the object of interest. Thus only one microscope objectiveis required for focusing both the two cameras as well as the excitationbeam onto the respective imaging plane or detection plane, respectively.Alternatively, only one camera can be used, in which embodiment thefocus control means are adapted for time-resolved focusing of the cameraon the at least two imaging planes, i.e. the at least two images of theobject of interest are subsequently acquired, and focusing of the camerais continuously (alternately) changed between the at least two imagingplanes. In the one-camera embodiment focus control means fortime-resolved focusing of the camera on the at least two imaging planes,in particular for alternately focusing of the camera on the at least twoimaging planes, are provided. These focus control means are thus onlylocated in the light path of the imaging system and not in light path ofthe Raman system.

The used camera(s) can be standard color CCD camera if the differentwavelength regions overlap with the different color regions (R,G or B)of the camera. It is also possible to use a CCD camera with speciallydesigned color sensitivity for different pixels, or to use a dichroicmirror and two monochrome cameras. Instead of CCD cameras also otherimaging devices can be used including: CMOS sensors (and photodiodearrays or scanning devices for other applications). For monochromaticOPSI a monochrome camera can be used.

The (relative) amount of defocus of the imaging system is preferablydetermined by use of a merit function. Examples for such merit functionsare the standard deviation of pixel intensity in the whole image or aregion of interest. When the system is focused sharp intensitytransitions are washed out and the standard deviation decreases. Otherexamples are the sum or average of gradient of pixel intensities in thewhole image or a region of interest. When the system is focused sharpintensity transitions are washed out and the gradient decreases. The(well known) Sobel gradient operator can calculate the gradient forexample, however, other gradient operators are also possible.

Common auto-focusing methods also use merit functions. However, the wayin which images at different depths are acquired according to thepresent invention is advantageous. As explained above, it is notpossible to quantify the amount of defocus in an absolute sense if theobject is not exactly known. In known auto-focusing methods the amountof focus is calculated for many settings of the system, and the bestfocus position corresponds to the position with the maximum merit value.Here, it is proposed to measure the amount of focus at two image planesat least. But not the absolute values are determined and evaluated, butonly a relative value, e.g. the difference between the amount for theimaging planes.

Different embodiments for the monitoring beam source are possible. It iswell possible to use only one single laser for emitting the monitoringbeam, or a single white light source. Further, appropriate filters canbe used to split up the generated white light beam into two or moremonitoring beams having different colors. According to a preferredembodiment two separate light sources, in particular light emittingdiodes (LEDs) are used emitting partial monitoring beams in differentwavelength areas, which beams are combined by a beam combination unitinto the monitoring beam for imaging the target region

It is preferred that the monitoring system is adapted for orthogonalpolarized a spectral imaging as mentioned above and as described inWO02/057759 A1.

The invention can not only be used in an analysis apparatus as describedabove, but relates also to an optical tracking system for continuouslytracking a point of a moving object, comprising a target system to befocused on the tracked point, a monitoring system, focusing means,defocus detection means and auto-focusing means as claimed in claim 12.The invention can be used in every system where an imaging system isused to locate and track a point, for example the focus of a laser beamor the detection volume of a spectroscopic system, continuously in 3dimensions at a specific position in a moving target. Examples include:(biomedical) laser surgery, laser cutting, laser welding, laser shaving,photodynamic therapy, remote sensing, and target and tracking inmilitary applications. Also the above described analysis apparatus couldbe regarded as including such an optical tracking system.

The invention will now be explained in more detail with reference to thedrawings in which

FIG. 1 shows a graphic representation of a first embodiment of ananalysis system according to the present invention,

FIG. 2 illustrates the principle used according to the presentinvention,

FIG. 3 illustrates the use of a merit function, and

FIG. 4 shows a graphic representation of a second embodiment of ananalysis system according to the present invention.

FIG. 1 is a graphic representation of a first embodiment of an analysissystem in accordance with the invention. The analysis system includes anoptical monitoring system (Iso) for forming an optical image of theobject (obj) to be examined. In the present example the object (obj) isa piece of skin of the forearm of the patient to be examined. Theanalysis system also includes a multi-photon, non-linear or elastic orinelastic scattering optical detection system (ods) for spectroscopicanalysis of light generated in the object (obj) by a multi-photon ornon-linear optical process. The example shown in FIG. 1 utilizes inparticular an inelastic Raman scattering detection system (dsy) in theform of a Raman spectroscopy device. The term optical encompasses notonly visible light, but also ultraviolet radiation and infrared,especially near-infrared radiation.

The monitoring system (lso) comprises a monitoring beam source (ls) foremitting a monitoring beam (irb) and an imaging system (img) for imagingthe target region, e.g. a blood vessel (V) in the upper dermis (D) ofthe patients forearm (obj). The monitoring beam source (ls) in thisexample comprises a white light source (las), a lense (l1) and aninterference filter (not shown) to produce light in the wavelengthregion of 560-570 nm. Further, a polarizer (p) for polarizing themonitoring beam (irb) is provided. The monitoring beam source (ls) isthus adapted for orthogonal polarized spectral imaging (OPSI).

In OPSI polarized light is projected by a microscope objective (mo)through a polarizing beam splitter (pbs) onto the skin (obj). Part ofthe light reflects directly from the surface (specular reflection).Another part penetrates into the skin where it scatters one or moretimes before it is absorbed or is re-emitted from the skin surface(diffuse reflection). In any of these scattering events the polarizationof the incident light is slightly changed. Light that is directlyreflected or penetrates only slightly into the skin will scatter onlyone or a few times before it is re-emitted, and will mostly retain itsinitial polarization. On the other hand, light that penetrates moredeeply into the skin undergoes multiple scattering events and iscompletely depolarized before re-emitted back towards the surface.

When looking at the object (obj) through a second polarizer or analyzer(A), oriented precisely orthogonal to that of the first polarizer (p),light reflected from the surface or the upper parts of the skin islargely suppressed, whereas light that has penetrated deep into the skinis mostly detected. As a result the image looks as if it wereback-illuminated. Because wavelengths below 590 nm are strongly absorbedby blood, the blood vessels appear dark in the OPSI image.

Generally, an image is obtained using a monochrome CCD camera. Bloodvessels are separated from other absorbing structures be means of size,shape and movement of blood cells. The imaging system (img) used in thepresent embodiment comprises an analyzer (a) mentioned above forallowing only light having a polarization orthogonal to the light of thepolarized monitoring beam (irb) to pass which is reflected back throughthe polarizing beam splitter (pbs) from the object (obj). Said light isfurther focused by a lens (l3) and split up by a beam splitter (bs) forreception by two CCD-cameras (CCD1, CCD2). These two cameras are focusedon two different imaging planes within the object (obj) which will beused for auto-focusing as will be explained below in detail.

The Raman spectroscopy device (ods) comprises an excitation system (exs)for emitting an excitation beam (exb) and a detection system (dsy) fordetection of Raman scattered signals from the target region. Theexcitation system (exs) can be constructed as a diode laser whichproduces the excitation beam in the form of an 785 nm infrared beam(exb). Of course other lasers can be used as excitation system as well.A system of mirrors and, for instance, a fiber conduct the excitationbeam (exb) to a dichroic mirror (f1) for conducting the excitation beam(exb) along the monitoring beam (irb) to the microscope objective (mo)for focusing both beams onto the object (obj).

The dichroic mirror (f1) also separates the return (monitoring) beamfrom scattered Raman signals. While the reflected monitoring beam istransmitted to the imaging system (img), elastically and inelasticallyscattered Raman light from the object is reflected at the dichroicmirror (f1) and conducted back along the light path of the excitationbeam. Inelastically scattered Raman light is then reflected by anappropriate filter (f2) and directed along the Raman detection path inthe detection system (dsy) to the input of a spectrometer with a CCDdetector. The spectrometer with the CCD detector is incorporated intothe detector system (dsy) which records the Raman spectrum forwavelengths that are smaller than approximately 1050 nm. The outputsignal of the spectrometer with the CCD detector represents the Ramanspectrum of the Raman scattered infrared light. In practice this Ramanspectrum occurs in the wavelength range beyond 800 nm, depending on theexcitation wavelength. The signal output of the CCD detector isconnected to a spectrum display unit, for example a workstation thatdisplays the recorded Raman spectrum (spct) on a monitor. Also acalculation unit (e.g. a workstation) is provided to analyze the Ramanspectrum and calculate the concentration of one or more analytes.

Regarding further details of the analysis apparatus in general and thefunction thereof reference is made to the above mentioned WO 02/057759A1.

The two cameras (CCD1, CCD2) are provided in the imaging system (img) toachieve continuous auto-focusing of the confocal Raman system (ods) in ablood vessel (V). This is required since patients will move during ablood analysis in lateral (z) as well as in transversal (x, y)directions. Therefore, continuous measurement and adjustment of theoptimal location of the confocal detection center is required.Transversal movements can be easily detected by the imaging system,whereas axial movements are much more difficult to detect. According tothe present invention the two cameras (CCD1, CCD2) thus take an OPSimage of the blood vessel (V); however, the imaging planes of the twocameras are slightly separated in axial direction (z). This is shown inmore detail in FIG. 2. While the first camera (CCD1) is focused onto afirst imaging plane (i1) above and in parallel to the detection plane(dp) onto which the Raman excitation and detection system (exs, dsy) arefocused, the second camera (CCD2) is focused onto a second imaging plane(i2) below and parallel to the detection plane (dp). The distance (d)between the two imaging planes (i1, i2) is in the order of or slightlylarger compared to the depth of field of the objective (mo), inparticular in the range from 0.5 to 20 μm. Preferably, the distances(d1, d2) of the imaging planes (i1, i2) from the detection plane (dp)are equal and fixed.

From the two images acquired by the two cameras (CCD1, CCD2) focusedonto the separate imaging planes (i1, i2) the amount of defocus ismeasured by a defocus detection means (ddm) with a merit function forboth cameras. Based on the determined difference in the amount ofdefocus and the known distances (d1, d2) between the imaging planes (i1,i2) and the detection plane (dp) the position of the microscopeobjective (mo) is adjusted by auto-focusing means (afm) such that theblood vessel (V) is imaged by both cameras (CCD1, CCD2) with an equalamount of defocus or, if that is beneficial, with a certain offset. Withthis information the Raman confocal detection volume can be continuouslylocated exactly inside the blood vessel (V).

A merit function is thus used to calculate a single number from animage. Depending on the type of merit function, this number is maximalor minimal for certain images. A merit function for automatic focusingis preferably chosen in such a way that it has its extreme value at thesharpest images. The extreme value of the merit function, however, isdifferent for different objects. Therefore, if the object is not exactlyknown, one cannot know if the image is properly focused from a singlemeasurement. Only by comparing the outcome of the merit function fordifferent settings of the system, one can determine the settings thatresult in the optimum focus. Shortly, one could say that if the objectis not exactly known a merit function only has relative and no absolutemeaning. Examples of a merit function are: Standard deviation of pixelintensity in the whole image or a region of interest. Sum of gradient ofpixel intensities in the whole image or a region of interest. The Sobelgradient operator can calculate the gradient for example. Other gradientoperators are also possible.

According to the present embodiment it is assumed that in the vicinityof the best focus position, the value of the merit function changessymmetrically above and below the best focus position. Therefore, if themerit function is calculated at two positions, and the outcome is equal,the optimum focus position must be located exactly in between thesepositions. If merit value 1 (M1; see FIG. 3) is larger compared to meritvalue 2 (M2) image plane 1 (i1) must be closer to the blood vesselcompared to image plane 2 (i2) and the central focus plane has to shiftupwards. Therefore, it is always known in which direction the focusposition has to be shifted from the relatives values of M1 and M2. Thisis important for a robust implementation of the system. The typicalshape of a merit function is illustrated in FIG. 3. In the currentembodiment it is just detected if M1 is larger or smaller compared toM2. Depending on the outcome the detection plane (dp) is moved up ordown by a fixed amount (e.g. 1 μm). This is the simplest embodiment. Ina more advanced embodiment it is also possible to determine the size ofthe step from the difference of defocus.

The complete image can be used for auto-focusing. However, differentblood vessels or parts thereof lie at different depths below the skinsurface. Therefore, it is more accurate to use a region of interestaround the best Raman measurement position for auto-focusing. In adifferent application with higher quality images, an accuracy of 1% ofthe depth of focus can be achieved with the method according to thepresent invention. Thus, for automatic focusing of the Raman excitationbeam, the acquired accuracy in the order of 1 μm can be obtained.

Another embodiment of an analysis apparatus according to the presentinvention is shown in FIG. 4. While the optical detection system (ods)is identical the imaging system (img). In the imaging system (img) onlyone CCD-camera (CCD1) is provided which, however, is adapted fortime-resolved reception of images from the target region at differentimaging planes. This can be achieved by continuously changing the focusof the camera (CCD1) between the first and the second imaging plane (i1,i2) by variation of the position of the lens (l2) controlled by focuscontrol means (fem). Thus, also with one camera the amount of defocusfor different imaging planes can be determined by use of a meritfunction as explained above so that finally also with this embodimentauto-focusing of the confocal Raman system in the blood vessel (V) canbe achieved.

Alternatives for continuously changing the focus of the camera (CCD1)are to move the camera or to use a switchable lens based on electrowetting.

In the above monochromatic OPSI embodiments are described having a whitelight source and a filter. Nevertheless, the invention can be applied inmany different embodiments including:

a) Monochrome OPSI embodiments having a single light source withwavelength below 590 nm, where the light source can be narrowband (LED,laser(diode)) or broadband (lamp), with a filter and having a monochromeimage module (CCD camera, CMOS sensor);

b) Bichromatic/multichromatic OPSI embodiments (in principle more thantwo colors can be used to extract extra information from the images)having

b1) a single light source with two colors, one above 600 and one below590 nm, which can be done by placing a filter with two transmissionregions in front of a white light source or by using a special lightsource that emits two different colors (e.g. a LED with two colors); itis also possible to use two or more light sources with different colorsand a beam combiner;

b2) a color camera or sensor, with different pixels sensitive to thecolors of the different light sources; this can be the normal RGBsensitivity curves or special curves (by special filters per set ofpixels); it is also possible to use a dichroic mirror and two monochromeimage sensors;

b3) a single broadband (white) light source and use one color camerawith special sensitivity curves corresponding to different wavelengthregions for the different pixels; alternatively, a dichroic mirror andtwo monochrome cameras with special filters in the optical path can beused;

b4) a single light source with a switchable spectrum or two lightsources that emit light in different wavelength regions that areswitched on alternately; a monochrome camera can then be used to capturethe images with different colors alternatively.

Compared to known auto-focusing techniques, the method according to thepresent invention has the advantage that the confocal detection volumecan be continuously located in the center of the blood vessel in whichthe blood shall be analyzed. The images that are compared are preferablymeasured simultaneously. Further, no moving elements are required in thetwo-camera embodiment, and a single high numerical aperture objectivecan be used for the monitoring system and the optical detection system.

1. An analysis apparatus, in particular a spectroscopic analysisapparatus, for analyzing an object comprising: an excitation system foremitting an excitation beam to excite a target region, a monitoringsystem comprising a monitoring beam source for emitting a monitoringbeam and an imaging system to image the target region, a detectionsystem for detecting scattered radiation from the target regiongenerated by the excitation beam, focusing means for focusing theimaging system on at least two substantially parallel imaging planes ata predetermined distance and for focusing the excitation system and thedetection system on a detection plane substantially parallel andin-between the imaging planes at predetermined distances defocusdetection means for determining the amount of defocus of the imagingsystem from the detection plane for the at least two imaging planes, andauto-focusing means for controlling the focusing means to commonlychange the focusing of the imaging system, the excitation system and thedetection system based on the determined amount of defocus and thepredetermined distances between the imaging planes and the detectionplane such that the difference between the determined amount of defocusfor the at least two imaging planes equals a predetermined amount.
 2. Ananalysis apparatus as claimed in claim 1, wherein the focusing means areadapted for focusing the imaging system on two imaging planes at equaldistances to the detection plane and wherein the auto-focusing means areadapted for controlling the focusing means such that the amount ofdefocus for each imaging plane is substantially equal.
 3. An analysisapparatus as claimed in claim 1, wherein the imaging system comprisestwo cameras each being focused on one of the imaging planes.
 4. Ananalysis apparatus as claimed in claim 1, wherein the imaging systemcomprises one camera and focus control means for time-resolved focusingof the camera on the at least two imaging planes, in particular foralternately focusing of the camera on the at least two imaging planes.5. An analysis apparatus as claimed in claim 1, wherein the defocusdetection means are adapted to determine the amount of defocus by use ofa merit function.
 6. An analysis apparatus as claimed in claim 1,wherein said merit function is the sum or average over all pixels in theimage or a region of interest of the intensity gradient, in particularas determined by the Sobel gradient operator.
 7. An analysis apparatusas claimed in claim 1, wherein the monitoring beam source comprises twolight sources for emitting partial monitoring beams in differentwavelength areas and a beam combination unit for combining the partialmonitoring beams into the monitoring beam.
 8. An analysis apparatus asclaimed in claim 1, wherein the monitoring beam source comprises asingle white light source having a filter for transmitting light in twoseparate wavelength regions and wherein the imaging system comprisesimaging means for color sensitive detection.
 9. An analysis apparatus asclaimed in claim 1, wherein the monitoring beam source comprises asingle light source with wavelength below 590 nm, said light sourcebeing a narrowband or broadband light source, a filter and a monochromeimaging system.
 10. An analysis apparatus as claimed in claim 1, whereinthe monitoring system is adapted for orthogonal polarized spectralimaging.
 11. An analysis method, in particular a spectroscopic analysismethod, for analyzing an object comprising the steps of: emitting anexcitation beam to excite a target region, emitting a monitoring beam toimage the target region by an imaging system, detecting scatteredradiation from the target region generated by the excitation beam,focusing the imaging system on at least two substantially parallelimaging planes at a predetermined distance, focusing the excitationsystem and the detection system on a detection plane substantiallyparallel and in-between the imaging planes at predetermined distances;determining the amount of defocus of the imaging system from thedetection plane for the at least two imaging planes, and controlling thefocusing to commonly change the focusing of the imaging system, theexcitation system and the detection system based on the determinedamount of defocus and the predetermined distances between the imagingplanes and the detection plane such that the difference between thedetermined amount of defocus for the at least two imaging planes equalsa predetermined amount.
 12. An optical tracking system for continuouslytracking a point of a moving object, comprising: a target system to befocused on the tracked point, a monitoring system comprising amonitoring beam source for emitting a monitoring beam and an imagingsystem to image the target region, focusing means for focusing theimaging system on at least two substantially parallel imaging planes ata predetermined distance and for focusing the target system on adetection plane substantially parallel and in-between the imaging planesat predetermined distances, defocus detection means for determining theamount of defocus of the imaging system from the detection plane for theat least two imaging planes, and auto-focusing means for controlling thefocusing means to commonly change the focusing of the imaging system andthe target system based on the determined amount of defocus and thepredetermined distances between the imaging planes and the detectionplane such that the difference between the determined amount of defocusfor the at least two imaging planes equals a predetermined amount. 13.An optical tracking system as claimed in claim 12, wherein said targetsystem comprises a light beam generation means for emitting a lightbeam, in particular a laser for emitting a laser beam, to be focused onthe tracked point of the object.
 14. An optical tracking system asclaimed in claim 12, adapted for use in the field of laser surgery,laser cutting, laser welding, laser shaving, photodynamic therapy, radiotherapy, remote sensing and target and tracking.